The present invention relates to a magnetic resonance imaging apparatus for generating a magnetic resonance signal (MR signal) by applying a gradient magnetic field pulse and high-frequency magnetic field pulse to an object to be examined which is placed in a homogeneous static magnetic field, and generating a magnetic resonance image (MR image) on the basis of this MR signal.
As is known well, the magnetic resonance phenomenon is a phenomenon in which when nuclei having unique magnetic moments are placed in a homogeneous static magnetic field, they resonantly absorb the energy of high-frequency magnetic fields that rotate at specific frequencies. A magnetic resonance diagnosis apparatus is designed to visualize the chemical, structural, microscopic information of a substance in a living body by using this magnetic resonance phenomenon.
Various types of visualizing techniques are available. The two-dimensional Fourier transform method (2DFT) is in the mainstream. In this 2DFT, a gradient magnetic field is used to give a magnetic resonance signal spatial positional information with a phase or frequency. This gradient magnetic field is superimposed on a static magnetic field having a very high strength of several kilogauss to 10 kilogauss (1 tesla). For this reason, the gradient coil spool considerably deforms at the leading and trailing edges of the gradient magnetic field, accompanying large noise. In a current fast imaging method represented by the echo planar method, a gradient magnetic field is alternated at high speed, the noise sometimes reaches 100 dB(A). This makes it obligatory for an object to wear earplugs or headphones.
To reduce such large noise, a proposal for suppressing the leakage of noise to the outside by housing a gradient coil spool in a sealed vessel has been disclosed in, for example, in Jpn. Pat. Appln. KOKAI Publication No. 63-246146, Jpn. Pat. Appln. KOKAI Publication No. 6-189932, U.S. Pat. No. 5,793,210, and Jpn. Pat. Appln. KOKAI Publication No. 10-118043 (Jpn. Pat. Appln. No. 8-274609). This sealed vessel is firmly formed by nonmagnetic aluminum or stainless steel to ensure sufficient pressure resistance.
Since these materials are conductive, they cause magnetic coupling with respect to leakage magnetic fields from the gradient coil. Obviously, if a so-called active shield gradient coil having a main coil surrounded by a shield coil is used, a leakage magnetic field can be sufficiently suppressed. However, it is inevitable that magnetic fields leak from the two ends of the main coil at which the shield coil ends. Owing to the above magnetic coupling, an eddy current flows in the sealed vessel. The sealed vessel itself then deforms because of this eddy current, which in turn causes noise.
According to these conventional sound insulating methods, although the air-born propagation of noise can be prevented by a vacuum state, the solid-born propagation of noise due to contact for fixing operation cannot be prevented. This solid-born propagation generates a large sound. That is, these methods have not provided satisfactory measures against noise. Although a method of suppressing solid-born propagation by supporting a gradient coil that produces large vibrations at a distance from the floor singly is available, if the position of the lower portion of a column supporting the gradient coil on the floor slightly shifts, the position of the gradient coil mounted on the upper portion of the column greatly shifts because the gradient coil is considerably spaced apart from the floor. This makes it difficult to adjust the position of the gradient coil, and a long period of time is required for installation/adjustment, resulting in a high cost. In addition, the means for supporting the gradient coil and the means for forming a sealed space around the gradient coil have complicated arrangements and require large numbers of parts. As a consequence, the overall apparatus has a complicated arrangement and the cost increases.
A technique is also known, which is designed to suppress the solid-born propagation of vibrations of the gradient coil itself by supporting the gradient coil through a vibration absorbing unit (damper).
At present, however, the noise reducing effects in these conventional magnetic resonance imaging apparatuses are concerned with only measures against noise originating from a gradient coil, and almost no measures are taken against noise caused by other portions, e.g., the vibrations of a cable connecting the gradient coil to an external power supply.
In a space (to be referred to as a bore) in which an object to be examined (patient) is inserted, the sound generated by the RF coil instead of the gradient coil is perceived as noise by the object. This sound makes the object feel unpleasant like noise from the gradient coil. According to the structure of a conventional RF coil, however, a conductive pattern is bonded to a spool and surrounded by a cover made of a hard resin or the like. No measures against the sound generated by the RF coil itself have been taken.
As described above, various silencing measures have recently taken. However, secondary problems have arisen owing to these silencing measures. As is known well, as one of RF coils, a whole body RF coil (to be referred to as a WB coil hereinafter) used for whole body imaging is available. This WB coil is generally placed closer to a patient than the remaining coils in the gantry. The WB coil includes a so-called transmitting/receiving coil for generating a spin excitation RF magnetic field and receiving an MR signal generated in the object and a so-called transmitting coil for only generating a spin excitation RF magnetic field, with reception being performed by another surface coil. In most magnetic resonance imaging apparatuses, a gradient coil for generating a gradient magnetic field is placed on the outer circumferential side of the WB coil. This gradient coil is used to apply, to an object to be examined, a magnetic field whose strength linearly changes depending on the position. For this purpose, the conductor of the gradient coil has a larger number of turns than that of the WB coil described above. In addition, the gradient coil must be switched with high efficiency in a frequency band much lower than the resonance frequency used in the MRI. That is, an energy loss that cannot be neglected tends to occur in the resonance frequency band.
For the WB coil (resonating at the resonance frequency in terms of a circuit), the electric loss caused by that gradient coil becomes a load that cannot be neglected, resulting in a deterioration in the generation efficiency of an excitation RF magnetic field or reception sensitivity.
Under the circumstances, to suppress the magnetic interference between the WB coil and the gradient coil, a technique of using a shield (a shield member such as copper foil) exhibiting a sufficiently low loss with respect to the resonance frequency is used (see U.S. Pat. No. 5,367,261). This shield is generally grounded.
On the other hand, it is known that when a shield is placed in this manner, the reception sensitivity of the WB coil is degraded by the shield (see, for example, “A Technique of Double Resonant Operation of F and H Quadrature Birdcage Coils” Magnetic Resonance in Medicine 19, 180-185 (1991)). However, the loss caused by the mutual interference between the WB coil and the gradient coil without any shield is higher than the loss caused by the WB coil with the shield. For this reason, in most MRI apparatuses, a shield is inevitably placed in the above manner. As is obvious from the above reference, as the distance between the shield and the WB coil increases, a deterioration in the efficiency of the WB coil can be suppressed.
As described above, a shield itself is used as the second best means. However, this shield can be easily fixed to a potential substantially different from that of the WB coil. So-called “detune” is known, which is a scheme of providing an electrical switch between a WB coil and a shield by using the above potential setting, and shifting the resonance frequency of the WB coil by ON/OFF-controlling the switch (see, for example, U.S. Pat. No. 5,053,711). By fixing the shield to, for example, zero potential, the shield can also be used as a zero potential surface for the WB coil and a circuit up to the WB coil.
The gradient coil mechanically vibrates due to an electromagnetic force acting when the coil is driven, and hence becomes a sound (noise) source. An apparatus designed to reduce this sound is also known, which has a structure in which a gradient coil is sealed in a cylindrical vessel whose internal pressure is controlled to a level substantially lower than the atmospheric pressure to suppress air-born sound propagation (see, for example, U.S. Pat. No. 5,793,210).
No suitable, specific proposal has not been made with regard to a case wherein the above shield is placed in such a silent type MRI apparatus.
If the conventional shield placement method is directly applied to a silent type MRI apparatus, the following arrangement can be expected. First of all, a WB coil is placed outside the inner circumferential wall member of the vessel, i.e., on the inner circumferential surface side of the vessel to place the coil as close to an object as possible so as to increase the S/N ratio. A shield is fixed to a given potential with respect to the WB coil, and is also used as zero potential for a switch and transmission system. The shield must therefore be placed outside the inner circumferential wall member of the vessel, i.e., on the inner circumferential side of the vessel. This is because, when the shield is to be used for these purposes, electric connection to the shield is preferably made near the WB coil. As a consequence, the WB coil, shield, vessel (in which the gradient coil is sealed), and static magnetic field magnet are arranged in the order named when viewed from the object in the radial direction of the magnet.
If, however, the shield is placed outside the inner circumferential wall member of the vessel, i.e., on the inner circumferential side of the vessel, in this manner, the distance between the WB coil and the shield in the radial direction of the magnet becomes smaller than that in the conventional structure without any vessel by the thickness of the vessel (the total thickness of the inner circumferential wall member and outer circumferential wall member) under the condition that the size of the overall gantry in the radial direction remains the same. As described above, this may increase the loss caused by the WB coil and cause a serious deterioration in function.
Furthermore, if the gradient coil is housed in the vacuum vessel in the above manner, since the discharge starting voltage in the vacuum vessel decreases, electric discharge tends to occur. At present, however, no measures have been taken against such electric discharge.